Reactive layer control system for prosthetic and orthotic devices

ABSTRACT

A variable gain impedance controller for use in a control system for controlling a motorized prosthetic or orthotic apparatus provided with a joint. The controller comprises a sensor input for receiving a signal indicative of an interaction between the apparatus and the ground, a torque sensor input for receiving a signal indicative of the torque at the joint, and a variable gain scheduler in communication with the sensor input to receive data therefrom thereby providing a variable torque gain. The variable gain impedance controller adjusts its control on the apparatus based on the variable torque gain and the indicated torque to increase the joint resistance to motion when the signal received from the sensor input indicates an interaction between the apparatus and the ground, and decrease the joint resistance to motion when the signal received from the sensor input indicates an absence of interaction between the apparatus and the ground.

CROSS-REFERENCE TO RELATED APPLICATION

The present application claims the benefits of U.S. provisional patentapplication No. 60/881,168 filed Jan. 19, 2007, which is herebyincorporated by reference.

TECHNICAL FIELD

The present invention relates to a reactive layer control system forprosthetic and orthotic devices.

BACKGROUND

Prosthetic and orthotic devices for restoring or replacing lostlower-limb functions have been available for many years. Until recently,both types of devices were found as purely mechanical linkages makingadvantageous usage of simple mechanisms in order to preclude kneebuckling in level walking stance phase, while still ensuring some formof swing motion during the aerial phase. While this type of device wasshown to be fairly efficient in restoring the structural aspects of thelower-limb role in gait, their incapacity to properly sustain the widevariety of lower-limb dynamics associated with the various gaitlocomotion activities performed on a daily basis appeared as asufficient limitation to sustain the development of more advanceddevices.

While significant efforts were directed towards designing more advancedmechanisms allowing easier adjustment, or more progressive action,through pneumatics and hydraulics, the rapid advances in energy storageand computer technologies soon allowed to extend the realm of capacitiesassociated with typical orthotic and prosthetic devices. Real-timeconfiguration of passive braking devices such as disclosed, for example,in U.S. Pat. No. 5,383,939 and US Patent Application Publication No.2006/0136072 A1, greatly improved the adaptability of prosthetic devicesto user gait specificities or to variations of the environment in whichthe locomotion tasks are performed. Moreover, these prosthetic devicesallowed the addressing of energy dissipative locomotion tasks in aphysiologically-compliant manner never seen before. Although showingincreased performance and dynamic adaptation with respect to thelocomotion tasks being undertaken when compared to their predecessors,this first generation of computer-controlled prosthetic devices stilllacked the adaptability and flexibility required to smoothly integrateinto users daily lives.

Integration of computer controls to the prosthetic and orthotic devicesbrought about the necessity for some sort of control system in order tolink sensory inputs to the now dynamically configurable actuator.However, the purely dissipative nature of these devices greatlysimplifies the problem as mechanical power exchanges between the userarid the device are unidirectional (i.e., user has to initiate all tasksand provide mechanical power).

Latest efforts in the field of advanced orthotic and prosthetic devices,such as disclosed, for example, in US Patent Application Publication No.2004/0181289 A1, partly resolved some of the limitations observed in thefirst generation of computer-controlled orthotic and prosthetic devicesby providing a fully motorized prosthetic platform, allowing to addressall major locomotion tasks, irrespective of their generative ordissipative nature. Requirements for computer-controlled system appearedquite more complex as the interactions between the user and theprosthetic or orthotic device were no longer solely initiated by theuser himself. Through the use of a two layer control system, themotorized prosthetic or orthotic device allowed to efficiently managethe mechanical power exchange between the user and the device, such thatthe synergy between user and motorized prosthetic or orthotic deviceglobally benefited the user. Adequate usage of the prosthetic ororthotic device capacity to generate mechanical power was observed tolead to increased gait quality and activity levels.

Nevertheless, the use of strict state machines to implement theartificial intelligence engine as the highest layer of the prosthetic ororthotic device control system is observed to impose a certain formalismon the manner in which the user executes typical locomotion tasks. Whilegenerating a certain learning burden on the user side, the use of firmtriggers in order to trigger either distinct state transition orspecific joint behavior greatly affects man-machine symbiosis. Moreover,limitations associated with the use of a strict state machine artificialintelligence engine when working in a highly variable environment (i.e.,external environment and user himself) are well known and quickly showup as robustness issues from a system perspective. Finally, processingassociated with the extraction of complex features associated withspecific locomotion task detection is also known to generate a latencybetween measurement of the sensors value and implementation of theactual actions, which is often observed to greatly affect the prostheticor orthotic device usability and performance.

Furthermore, common prosthetic or orthotic devices lack the ability toproperly reproduce natural knee joint behavior and dynamic propertieswhen used in a context that significantly differs from typicallocomotion tasks. While generation of proper joint dynamics duringcyclical locomotion portions ensure high symbiosis and user benefits,limitations observed in the capacity to reproduce natural jointcompliance, or motions, in either non-locomotor or non-cyclical taskssignificantly affect orthotic, or prosthetic, device usability and,accordingly, associated user benefits.

Based on these last observations, it clearly appears that requirementsfor an improved orthotic and prosthetic control system exist. Morespecifically, a need to develop a control system architecture andassociated engines that are able to sustain more efficiently limitedambulation, as well as non-cyclical and cyclical gait for userssuffering of either amputation of the lower-limb or dysfunctionrequiring the use of an orthosis or prosthesis exists.

SUMMARY

In accordance with an aspect of the present invention there is provideda variable gain impedance controller for use in a control system forcontrolling a prosthetic or orthotic apparatus provided with a joint,the controller comprising:

-   -   a sensor input for receiving a signal indicative of an        interaction between the apparatus and the ground;    -   a torque sensor input for receiving a signal indicative of the        torque at the joint and    -   a variable gain scheduler in communication with the sensor input        so as to receive data therefrom thereby providing a variable        torque gain;    -   wherein the variable gain Impedance controller adjusts its        control on the apparatus based on the variable torque gain and        the indicated torque so as to increase the joint resistance to        motion when the signal received from the sensor input indicates        an interaction between the apparatus and the ground.

In accordance with another aspect of the present invention there isprovided a variable gain impedance controller for use in a controlsystem for controlling a motorized prosthetic or orthotic apparatusprovided with a joint, the controller comprising:

-   -   a sensor input for receiving a signal indicative of an        Interaction between the apparatus and the ground;    -   a torque sensor input for receiving a signal indicative of the        torque at the joint; and    -   a variable gain scheduler in communication with the sensor input        so as to receive data therefrom thereby providing a variable        torque gain;    -   wherein the variable gain impedance controller adjusts its        control on the apparatus based on the variable torque gain and        the indicated torque so as to decrease the joint resistance to        motion when the signal received from the sensor input indicates        an absence of interaction between the apparatus and the ground.

In accordance with a further aspect of the present invention there isprovided a variable gain impedance controller for use in a controlsystem for controlling a motorized prosthetic or orthotic apparatusprovided with a joint, the controller comprising:

-   -   a sensor input for receiving a signal indicative of an        interaction between the apparatus and the ground;    -   a torque sensor input for receiving a signal indicative of the        torque at the joint; and    -   a variable gain scheduler in communication with the sensor input        so as to receive data therefrom thereby providing a variable        torque gain;    -   wherein the variable gain impedance controller adjusts its        control on the apparatus based on the variable torque gain and        the indicated torque so as to a) increase the joint resistance        to motion when the signal received from the sensor input        indicates an interaction between the apparatus and the ground,        and b) decrease the joint resistance to motion when the signal        received from the sensor input indicates an absence of        interaction between the apparatus and the ground.

BRIEF DESCRIPTION OF THE FIGURES

Embodiments of the invention will be described by way of example onlywith reference to the accompanying drawings, in which:

FIG. 1 is a block diagram of the interaction between various controlsystem layers and major building blocks of a motorized prosthetic and/ororthotic device;

FIG. 2 is an isometric view of a motorized knee prosthesis;

FIG. 3 is a schematic representation of the lower-limb mechanical powerexchange during ground contact phase;

FIG. 4 is a schematic representation of the lower-limb mechanical powerexchange during aerial phase;

FIG. 5 is a block diagram of a variable gains impedance controller basicformulation;

FIG. 6 is a flow diagram of a gain scheduling mechanism and associatedreference engine;

FIG. 7 is a block diagram of a variable gains impedance controller withbreaking feedback transfer function;

FIG. 8 is a chart illustrating the braking reactive behavior activationsubspace; and

FIG. 9 is a block diagram of a variable gains impedance controller withenergy injection feedforward function.

DETAILED DESCRIPTION

Generally stated, the non-limitative illustrative embodiment of thepresent invention provides a reactive layer control system for motorizedprosthetic or orthotic devices for restoring lost locomotor functions,or facilitate gait re-education resulting from various pathologiesoccurrence. The reactive layer control system is part of a multi-layeredcontroller and is based on Impedance control, which directly manages asubset of lower-limb joint behaviors allowing the sustaining of highlyefficient mechanical power exchanges between the user and a prostheticor orthotic apparatus.

Referring to FIG. 1, there is shown a block diagram of a motorizedprosthetic and/or orthotic device 200 which comprises a multi-layeredcontroller 100 that may be used to control a motorized prosthetic ororthotic apparatus 140 such as, for example, the motorized kneeprosthesis 10 of FIG. 2.

Referring now to FIG. 2, the motorized knee prosthesis 10 includes aproximal connector 17 sitting on top of an actuator 12 which is axiallymounted at the knee joint 11 level. In this example, the actuator 12 maybe, for example, a DC brushless motor serially connected to a reductionmechanism. The reduction mechanism of the actuator 12 allows theconversion of the motor high-speed low-torque output characteristicsinto a low-speed high-torque output that is more coherent with therequirements associated with the human knee joint role in most commonlyencountered locomotor tasks. A second transmission stage is thenprovided in order to connect the reduction mechanism output to the shankstructure 13 of the motorized knee prosthesis 10. This secondtransmission stage is composed of a compliant linkage 14, allowing bothmeasurement of the net torque present at the interface between the shankstructure 13 and the actuator 12 output and high-efficiency levelwalking stance flexion energy storage and return.

The motorized knee prosthesis 10 also integrates sensors required tosustain the multi-layered controller 100 (see FIG. 1). A first positionencoder (not shown) is integrated to the transmission assembly of theactuator 12 such that the relative position between the user thighsegment (not shown) and the reduction mechanism output is measured inreal-time. Net torque present at the interface between the shankstructure 13 and the actuator 12 output is measured through thedeflection of the compliant linkage 14 transmitting motion between bothparts 12, 13, using a second position encoder (not shown) mounted in thetransmission assembly of the actuator 12 for that purpose. A load cellassembly 19 containing one or two load cells 16 is located at the distalshank portion 15, between the shank structure 13 and the distalconnector 18 of the motorized knee prosthesis 10, to quantify the stressfound in the distal shank portion 15.

It is to be understood that although the motorized knee prosthesis 10described above has been given as an example of the motorized prostheticor orthotic apparatus 140, the multi-layered controller 100 may besimilarly used with other motorized prostheses or orthoses havinggeneral characteristics similar to that of the motorized knee prosthesis10. More specifically, the multi-layered controller 100 may be similarlyused with motorized or actuated prostheses or orthoses having means formeasuring the net torque of its actuator output, means for detectingground contact and means for measuring the position of its actuator.

Referring back to FIG. 1, the multi-layered controller 100 isparticularly well suited for optimizing the synergy between a user andmotorized prosthetic and/or orthotic device 200 through theimplementation of motorized prosthetic or orthotic apparatus 140 jointbehaviors similar to those which may be observed on a sound human kneejoint.

The multi-layered controller 100 includes, but is not limited to, threelayers herein referred to as the learning layer 110, the inference layer120 and the reactive layer 130. Layering of the multi-layered controller100 aims at providing a systematic way of distributing thefunctionalities of the multi-layered controller 100 with respect totheir level of abstraction, hence allowing the definition of a coherentand straightforward architecture. It is to be understood that themulti-layered controller 100 may include more or less than three layers.

In order to interact with the environment 150 it evolves in, themotorized prosthetic and/or orthotic device 200 includes, but is riotlimited to, sensors 142 providing information about the environment 150and the motorized prosthetic or orthotic apparatus 140 to themulti-layered controller 100, and one or more actuator 144, controlledby the multi-layered controller 100, to generate behavior allowing tosustain an optimal interaction with the environment 150. For example, inthe case of the motorized knee prosthesis 10 of FIG. 2, the first andsecond position encoders (not shown), the compliant linkage 14 and theload cells 16 would compose sensors 142 while the actuator 12 wouldcompose actuator 144.

Multi-Layered Controller

While all three layers 110, 120, 130 of the multi-layered controller 100operate as stand-alone entities, information is propagated across thelayers 110, 120, 130 such that lower-level layer mechanisms maybeneficiate from information provided by higher-level layers. In such amulti-layered controller 100, decisions are performed independentlyinside of the different layers 110, 120, 130 characterized by differentdata abstraction levels, while propagation of information towards thelower-level layers ensures the adaptation of the lower-level layermechanisms. In a similar fashion, information provided by thelower-level layers is merged into higher abstraction levelrepresentations when moved towards the higher-level layers.

Learning Layer

The learning layer 110 represents the highest data abstraction level ofthe multi-layered controller 100. More specifically, the dataabstraction level associated with this layer is characterized as theuser data. Functionalities associated with this level of themulti-layered controller 100 relate to the recursive improvement of thehigh level strategies to address locomotion tasks, as they areaccomplished, and their relative performance assessed. At this level,representations of the user gait specificities identified during theevolution of the synergy between the user and the motorized prostheticand/or orthotic device 200 are updated and stored.

Inference Layer

The inference layer 120 contains locomotion task level information andfunctionalities. At this abstraction level are found the enginesrequired to perform locomotion task identification and characterization.Most of the work performed at this level consists in extracting typicalfeatures from the raw input data stream from the sensors 142 such thatthe locomotion task performed by the user may be characterized andsystem behavior adjusted according to the high-level information readilyavailable from the learning layer 110.

Reactive Layer

At the lowest level, the reactive layer 130 sustains the implementationof general classes of joint behaviors that are common to a large subsetof locomotor and non-locomotor activities. Similarly to the arc-reflexpresent in the human locomotor system, the reactive layer 130 is used inorder to directly link low-level sensory inputs from the sensors 142 toeither joint actions or behaviors of motorized prosthetic or orthoticapparatus 140 through the actuator(s) 144. Major benefits associatedwith integration of such reactive behaviors in a multi-layeredcontroller 100 arise from the fact that these behaviors allow a reduceddependency on high-level decisions in order to implement specificactions.

Reducing dependency between high-level decision making and actionsallows the reduction of latencies related to processing of high-levelinformation and to generate simpler, more robust, mapping betweensensory inputs from the sensors 142 and actions via the actuator(s) 144.Moreover, while generating more human-like behaviors from a userperspective, such implementation provides greater flexibility to theuser who now find himself in full control of the motorized prosthetic ororthotic device's 200 most basic behaviors.

Linking low-level triggering mechanisms to the basic joint behaviorsincreases system conviviality and realm of performance, as it is stillpossible to trigger higher-level mechanisms generating more complexjoint behaviors or motions, that will be simply defined asspecialization of the more basic behaviors. This way, complex motions orelaborate joint behaviors may be generated from adding specificinformation to the basic behavior implicitly provided by thelowest-level layers of the multi-layered controller 100.

An example of a controller implementing a learning layer 110 and aninference layer 120 is shown in US Patent Application Publication No.2006/0122710 A1 entitled “CONTROL DEVICE AND SYSTEM FOR CONTROLLING ANACTUATED PROSTHESIS” by Bedard. The reactive layer 130 will be furtherexplained below.

Reactive Layer Control System

A reactive layer control system for motorized prosthetic or orthoticdevices according to an illustrative embodiment of the present inventionrelates to the definition of a reactive layer engine which may be usedwithin the context of a multi-layered controller, such as themulti-layered controller 100 of FIG. 1.

The reactive layer control system is based on a variable gain impedancecontroller and has for goal to increase the synergy between the user andthe motorized prosthetic and/or orthotic device 200 for all types oflocomotion activities while directing specific attention towards systemperformance improvement for non-cyclical ambulation tasks. Improvementof motorized prosthetic and/or orthotic device 200 performance forlimited ambulation locomotion tasks requires a greater flexibility ofthe reactive layer 130 such that general motorized prosthetic and/ororthotic device 200 behaviors may fulfill user requirements in anon-model based framework. Use of a model-based framework to managelocomotion tasks not presenting obvious physiological characteristics orhigh inter-subject variability presents severe limitation to themotorized prosthetic and/or orthotic device 200. Failure to generate acomplete and robust mapping between the sensory inputs and the requiredactions actually impairs the general feasibility of a model-basedframework.

However, definition of basic motorized prosthetic or orthotic apparatus140 joint behaviors showing high correlation to the lower-limb jointsphysiological behavior and their integration to the lowest level of amulti-layered controller, such as the multi-layered controller 100 ofFIG. 1, allows to implicitly fulfill specific tasks requirements, whileleaving full control of the motorized prosthetic and/or orthotic device200 behavior to the user.

The overall objective of the reactive layer control system is to reducethe dependency between decision and action for a general class ofbehaviors that may be compared to human arc-reflex. The general class ofbehaviors is found as the basic behaviors underlying most of thelocomotion tasks. Implementation of reactive behaviors in the motorizedprosthetic and/or orthotic device 200 leads to an increase in robustnessand a significant reduction of the constraints associated withtraditional decision process for a system where all actions aresustained by explicit decisions.

High fidelity reproduction of the human knee joint natural behavior isrequired in order to properly sustain limited ambulation tasks,generally improve mechanical power exchange management and easeconstraints related to synchronization of the motorized prosthetic ororthotic apparatus 140 joint behavior transition with overall dynamicsof the user.

Human knee joint role in gait for locomotor and non-locomotor tasks maybe classified in general classes of joint behaviors as illustrated inthe following table:

TABLE 1 Joint behavior classes Reactive Controller Joint behavior classBehavior Behavior Passive motion without force perturbation force (e.g.,aerial phase) matching Isometric support without motion perturbationforce (e.g., contact phase) rejection Eccentric energy dissipationbraking Concentric mechanical power generation energy injection

These general classes of joint behavior may then be directly managedthrough the Implementation of an associated reactive layer controllerbehavior.

Impedance Control

The reactive layer control system is built around a typicalimplementation of an impedance controller. The impedance controller wasfirst introduced by Hogan in 1985. see [1], [1]and [3], as a first stepin defining a general and unified approach to the control ofmanipulation by robotic devices. While being very general, this specificcontrol scheme is rather well suited for managing tasks where highlydynamic interactions between a robotic device and the environment arepresent. Apart from other traditional control schemes targeting theindividual control of actuator variables such as torque or positioncontrol, impedance control implements a scheme where the overallobjective is defined as implementing a dynamic relationship betweenactuator variables, such as torque and position. In other words, theimpedance controller does not try to track specific trajectories, butinstead attempts to regulate the relationship between actuator velocityand force. The dynamic relationship between actuator force and velocityis generally known as “mechanical impedance”. This nomenclature arisefrom similarity to the electrical quantity found as the ratio of aneffort variable (i.e. voltage) to a flow variable (i.e. current). In theLaplace domain, mechanical impedance may be represented as follows:

$\begin{matrix}{{Z(s)} = \frac{F(s)}{V(s)}} & {{Equation}\mspace{14mu} 1}\end{matrix}$

-   -   where        -   Z(s) is the mechanical impedance;        -   F(s) is the actuator force; and        -   V(s) is the actuator velocity.

At the opposite, mechanical admittance describes the dynamicrelationship between actuator velocity and force. In the Laplace domain,mechanical admittance may be represented as follows:

$\begin{matrix}{{Y(s)} = \frac{V(s)}{F(s)}} & {{Equation}\mspace{14mu} 2}\end{matrix}$

-   -   where        -   Y(s) is the mechanical admittance;        -   V(s) is the actuator velocity; and        -   F(s) is the actuator force.

While the relationships represented by Equations 1 and 2 are generallyinterchangeable for linear systems operating at finite frequencies, thisis not the case for typical prosthetic or orthotic applications, whichare generally highly non-linear. Moreover, due to the input-outputspecificities of the mechanical system behaviors described above, it isonly possible to physically connect components of different nature.Failure to fulfill this requirement actually makes impossible propermanagement of the mechanical power exchanges at the interface ports, asboth components will try to impose the same physical quantity.

As far as the description of lower-limb joints physical behavior isconcerned, one has first to consider that the structure of the humanlower-limb, coupled with locomotor and non-locomotor gait specificities,generate two different mechanical configurations 30, 40, representedconceptually in FIGS. 3 and 4, respectively. In a first configuration30, the lower-limb joint 34 is located between the external environment(i.e. ground) 32 and the user's upper body mass 36. In a secondconfiguration 40, the lower-limb joint 44 is located between the hangingdistal limb mass 42 and the user's upper body mass 46, and is thussubmitted to significant dynamic efforts. For the first configuration30, velocity constraints are imposed on the lower-limb joints 34 (i.e.healthy joints and prosthetic or orthotic joints) by the ground 32 onthe distal end and by the user's upper body mass 36 on the proximal end.As for the second configuration 40, velocity constraints are imposed onthe lower-limb joints 44 (i.e. healthy joints and prosthetic or orthoticjoints) by the distal limb 42 (residual limb or prosthesis) dynamics onthe distal end and by the user's upper body mass 46 on the proximal end.

It is to be understood that “ground” is meant to mean, in the context ofthis specification, any surface on which a user may use the motorizedprosthetic and/or orthotic device 200 during locomotion activities

Ground Contact Phase

FIG. 3 provides a high-level representation of the lower-limb componentsinteractions during ground contact phase through the use of mechanicalimpedance/admittance. An impedance is a system characterized by itscapacity to accept a flow input V(s) (i.e., velocity) and yield aneffort F(s) (i.e., force). An admittance is a system characterized byits capacity to accept effort inputs F(s) (i.e., force) and yield a flowV(s) (i.e., velocity).

In order for mechanical power exchange to take place between both typesof system, input-output variables V(s) and F(s) must be matched. Sinceit is not possible to impose a velocity to the ground 32, it is modeledas an admittance. Connecting any type of lower-limb device to the ground32 then requires this latter to be defined as an impedance. Furthermore,the upper body mass 36 is also modeled as a admittance as it may onlyimpose velocity on the lower-limb joints 34 and segments. Force observedin the lower-limb joints 34 during the ground contact phase then arisefrom the impedance of the joints themselves. Thus, it may be observedthat in configuration 30, the lower-limb joints 34 form a systemoptimally represented as an impedance interacting with the user's bodymass 36 and ground 32, both modeled as admittance.

Aerial Phase

FIG. 4 provides a high-level representation of the lower-limb componentsinteractions during the aerial phase. In this configuration 40, thelower-limb joints 44 are mainly submitted to the effects of the distallimbs mass 42 and upper body mass 46. Again, a mass being characterizedas an element that accepts force F(s) as input while yielding velocityV(s) as output, it appears necessary to define the behavior of thelower-limb joints 44 as an impedance in order to ensure that stablemechanical power exchanges may take place. Based on these observations,it appears clear that the definition of any lower-limb prosthetic ororthotic devices, motorized or not, must take the form of an impedanceif it is desired to optimize user-device synergy and properly managemechanical power exchange

Furthermore, this is also coherent with the role of the lower-limbjoints in cyclical locomotion activities, which consists in absorbingshocks generated by the ground contact occurrence, such that body centreof mass trajectory is regulated and smooth progression occurs. Use of animpedance controller in order to manage the prosthetic or orthotic jointbehavior then appears as a straightforward solution to the problem athand.

Impedance Controller

As previously introduced, the impedance controller differs from moretraditional motion control schemes through the fact that it does notattempt to track specific control variables, such as force or position,but implements a scheme that allows regulation of the actuator 144 (seeFIG. 1) output mechanical impedance. Furthermore, this specific schemeimplicitly manages transitions where the actuator 144 physicalconfiguration is changing from non-interacting configuration 40 with theenvironment to an interacting configuration 30 (see FIGS. 3 and 4),which is not the case with other types of control schemes.

Referring to FIG. 5, there is shown a basic formulation of a variablegains impedance controller 50 that may be implemented at the reactivelayer 130 (see FIG. 1). The motorized prosthetic or orthotic apparatus140 under control is represented by a Laplace-domain double integrator52. First, position Θ and velocity {dot over (Θ)} feedback loops 51, 53are closed to form a tracking controller where position Θ_(d) andvelocity {dot over (Θ)}_(d) set-points are used as comparison values forthe feedback position Θ and velocity {dot over (Θ)} values. Furthermore,variable gains K_(P) and K_(D) are applied to both position and velocityerror terms (i.e., difference between the set-point values Θ_(d) and{dot over (Θ)}_(d), and the measured feedback values Θ and {dot over(Θ)}). {umlaut over (Θ)}_(d) represents the acceleration set-point.

Additionally to what would otherwise be considered as a simpleproportional-derivative position controller, interaction between theactuator 144 output port position Θ, with the position perturbationcreated by the environment Θ_(e), generates a generalized force τ_(A)quantifying the interaction force between the actuator 144 output andits environment. This measured force value τ_(A) is then used as anegative feedback loop 55, creating an actuator 144 set-point value ofthe same amplitude as the interaction force, assuming unitary forcefeedback gain K_(A), but of opposite sign. Assuming that satisfactoryforce sensing capacities are available, such system would then show aninfinite impedance (i.e. any perturbation force applied on the actuator144 output would be immediately converted to an opposite actuator 144reaction, leading to no displacement of the actuator 144 under theaction of the external force) without any contribution of the position Θand velocity {dot over (Θ)} terms. Modification of the force feedbackterm gain K_(A) allows the scaling down of the actuator 144 mechanicalImpedance by reducing the amount of force that is sent back as actuator144 set-point.

In such a variable gains impedance controller 50, position Θ andvelocity {dot over (Θ)} terms are used to generate the system dynamicresponse to either effects of external perturbation Θ_(e) ormodifications to the system's position Θ_(d) and velocity {dot over(Θ)}_(d) set-points. Such combination of proportional-derivativeposition control and the measured interaction force allows the fullcompensation of any perturbation present at the system mechanicalinteraction port, while still allowing to enforce a specific dynamicresponse.

A final gain, the mass gain M_(d) ⁻¹, affects the complete actuator 144force set-point and is generally considered to allow simulation ofsystem apparent inertia through appropriate scaling of the variablegains impedance controller 50 output. While the variable gains impedancecontroller 50 basic behavior described above already provides aninteresting framework for managing interactions and mechanical powerexchanges between the user and the motorized prosthetic and/or orthoticdevice 200, coupling of the variable gains impedance controller 50 witha gain scheduling mechanism, which will be described further below, isshown to further extend the realm of implicitly supported behaviors.While use of high-level engines to manage gain scheduling allows theadaptation of prosthetic or orthotic apparatus 140 joint behaviors basedon the nature of the locomotion tasks currently executed, lower-levelgain scheduling engines allow the adaptation of the variable impedancecontroller parameters such that optimal use of the inherent behaviors ofthe variable gains impedance controller is made without compromisingsystem performance from an user standpoint.

The above described variable gains impedance controller 50 may be usedto implicitly implement the first two joint behavior classes of Table 1,namely the Passive and isometric classes, while its general structuremay be used to explicitly integrate the third and fourth joint behaviorclasses, namely the Eccentric and Concentric classes.

Force Matching and Force Rejection Implementations

As discussed above, the first two joint behavior classes, i.e. Passiveand isometric, are addressed through proper usage of the implicitbehaviors of the variable gains impedance controller 50. These first twojoint behaviors classes are considered the most basic ones as alllocomotion task will first be characterized as being composed of one, orboth, of these behaviors.

The behavior of the Isometric joint behavior class corresponds to ajoint behavior where force without motion is generated, and will beherein associated to a joint behavior where it is desired to providestability and support, without generating any motion. This behavior isassociated with the stance phase of all cyclical and non-cyclicallocomotion tasks, where it is advantageous from a safety and usabilitystandpoint to be able to provide support to the user without enforcingany motion.

Referring back to FIG. 5, from a variable gains impedance controller 50standpoint, such behavior corresponds to an infinite impedance of theactuator 144 output with respect to the effects of externalperturbations Θ_(e). As previously introduced, such behavior isimplicitly generated by the variable gains impedance controller 50assuming that the force feedback gain K_(A) is adequately selected. Inorder for an infinite impedance behavior to take place, magnitude of themeasured interaction force must be very similar to the one of the forceactually imposed on the actuator 144 output, force losses in theactuator 144 and transmission must be accounted for and latency of theactuator 144 reaction with respect to the external perturbation Θ_(e)must be small enough not affect the closed-loop stability of thevariable gains impedance controller 50.

With reference to the motorized knee prosthesis 10 of FIG. 2, thevariable gains impedance controller 50 force feedback loop 55 valueτ_(A) may be provided by the measurement of the net torque found at theinterface between the actuator 12 output and shank structure 13. Aspreviously introduced, measurement of the deflection of the compliantelement 14 provides a direct measure of the net torque. While measuringthe net torque through a compliant element 14 greatly reduces thesensing bandwidth with respect to other technologies, such technique isshown to provide satisfactory results in the context where human motionsare showing only limited bandwidth and allow some flexibility withrespect to system reaction latency.

From a usability perspective, it is advantageous for the motorizedprosthetic and/or orthotic device 200 (see FIG. 1) to provide support tothe user without actually impairing his capacity to voluntarily move thesystem 200 from a given position to another one, while maintainingprosthetic or orthotic apparatus 140 joint stability. Such setting Isfound through the adequate adjustment of the force feedback gain K_(A)until satisfactory joint impedance is obtained, i.e. leading to anon-infinite joint impedance, with respect to user ability level andpersonal preferences.

In a similar manner, the passive joint behavior may be directlyimplemented using the inherent characteristics of the variable gainsimpedance controller 50. As the Passive joint behavior class is directlyassociated with the aerial phase of any locomotion task, it isadvantageous to make the motorized prosthetic and/or orthotic device 200as compliant as possible, such that overall user-device synergy maybenefit from the direct interactions between user residual limb motionsand the inertial properties of the motorized prosthetic or orthoticdevice 200. Moreover, making the motorized prosthetic and/or orthoticdevice 200 as compliant as possible during the aerial phase allows theminimization of the inertia reflected at the stump-socket interface (forexample, the socket, which is not shown, connects to the proximalconnector 17 of the motorized knee prosthesis 10 of FIG. 2). This way, asignificant reduction of the apparent weight of the motorized prostheticand/or orthotic device 200 is obtained from a user perspective, whilethe motorized prosthetic and/or orthotic device 200 also becomes easierto manipulate.

From a variable gains impedance controller 50 standpoint, generating aminimum impedance behavior during the aerial phase requires the actuator144 command signal to act in such a way that the force measured at theactuator 144 output remains null or negligible. Obviously, this requiresthe actuator 144 output to move in the same direction as the shankstructure 13, such that the net force between both parties remains nullor negligible. Assuming again a null contribution of the proportionaland derivative terms of the variable gains impedance controller 50,i.e., K_(P)≅0 and K_(D)≅0, this behavior is achieved by modifying theforce feedback gain K_(A) value such that the measured interaction forcenow becomes a positive set-point to the actuator 144, i.e., achieved byinverting the sign of the force feedback gain K_(A).

Assuming proper selection of the force feedback gain K_(A) value andminimal latency of the actuator 144 command with respect to the measuredforce, minimal joint impedance is obtained. Such scheme also providesthe benefit of compensating for the actuator 144 mechanicalnon-linearities, which are known to greatly affect the passive dynamicproperties of motorized prosthetic or orthotic systems. This is themajor difference between using null gains in a position control schemeand performing perturbation force matching with the variable gainsimpedance controller 50. While the position control system would simplyturn off the actuator 144, the variable gains impedance controller 50with the perturbation force matching approach allows to compensate foractuator 144 dynamic non-linearities, i.e. transmission back-drivingtorque, actuator motor cogging, actuator motor and transmission bearingsfriction, hence really minimizing joint impedance. In fact, in themotorized knee prosthesis 10 of FIG. 2, only the friction found in theshank structure 13 bearings is not compensated through the perturbationforce matching scheme.

Full compensation of the actuator 144 dynamic non-linearities wouldrequire measurement of the external perturbation Θ_(e) force at anotherlevel of the structure, for example at the foot-ground interface.Nevertheless, measurement of the external perturbation Θ_(e) force atthe actuator 144 output is found more flexible with respect tolower-limb mechanical configuration and ensure high co-linearity betweenforce measurement and actuator 144 output.

As introduced earlier, modification of the gains of the variable gainsimpedance controller 50 is required in order to change the jointbehavior of the motorized prosthetic or orthotic apparatus 140 from afinite impedance level to a null impedance level. This change is limitedin scope and is directly correlated with the lower-limb mechanicalconfigurations 30, 40, represented conceptually in FIGS. 3 and 4, i.e.,ground contact and aerial phase respectively. In order for the reactivelayer 130 behavior to take place without affecting overall motorizedprosthetic and/or orthotic device 200 performance, it is advantageous tominimize decisional overhead and device behavior transition latency froma user perspective.

Gain Scheduling Mechanism

Referring to FIG. 6, there is shown a simplified block diagram of alow-level gain scheduling mechanism and associated inference engine 60that may be used in order to modify the behavior of the motorizedprosthetic and/or orthotic device 200 where a transition from thelower-limb interacting mechanical configuration 30 to thenon-interacting configuration 40, or the opposite, is detected. First,raw sensor signals 61 from the sensors 142 (see FIG. 1) are provided toa detection mechanism in the form of an inference engine 62 in order toidentify if the lower-limb mechanical configuration is interacting 30 ornon-interacting 40 (see FIGS. 3 and 4) by, for example, detecting groundcontact. Various types of sensors 142 may be used in order to sustainthe decisional process of the inference engine 62, for exampleinstrumented plantar orthosis, accelerometers, digital switches, loadcells, etc. Advantageously, with reference to FIG. 2, a load cellassembly 19 containing one or two load cells 16 located at the distalshank portion 15 maybe used to provide the raw sensor signals 61.

The decisional process of the inference engine 62 may implement low-passfiltering of the raw sensor signals 61 combined with single valuehysteretic thresholding of the low-pass filtered raw sensor signals 61in order to identify the lower-limb mechanical configuration 30, 40.Based on the result of the thresholding process, a perturbation forcematching 64 or perturbation force rejection 66 gain scheme is providedto the dynamic gain update process 68.

The dynamic gain update process 68 then proceeds to the dynamic updateof the gains of the variable gains impedance controller 50 using, forexample, linear transition patterns or other patterns, where thetransition duration is configurable in order to adapt to user personalpreferences and gait specificities. In the illustrative embodiment, onlythe proportional K_(P), derivative K_(D), and force feedback K_(A) gainsare modified. The mass gain M_(d) ⁻¹ is maintained unitary and constant.Moreover, while the force feedback gain K_(A) transition from a negativevalue to a positive value upon occurrence of a ground contact event, theproportional K_(P) and derivative K_(D) gains are maintained to the samevalues, which are voluntarily selected close to zero. Based on resultsfrom experimental trials, a substantially unitary positive forcefeedback gain K_(A) during the ground phase coupled to a substantiallyunitary negative feedback gain K_(A) during the aerial phase leads to anoptimal gain configuration.

Reactive implementation of the Passive and Isometric joint behaviorclasses by the variable gains impedance controller 50 provides theunderlying foundations to the implementation of any locomotion task andwill also define the default behavior of the motorized prosthetic and/ororthotic device 200. Based on the fact that the combination of thesebehaviors will sustain all limited ambulation tasks, while leaving theuser in full control of the management of mechanical power exchanges,benefits arising from such a scheme are multiple, namely:

-   -   no requirement for a orthotic or prosthetic device-user        synchronization mechanism as transitions are initiated by the        user and the reaction time of the motorized prosthetic and/or        orthotic device 200 is quite short;    -   no requirement for high-level detection of transitions between        isometric and passive joint behavior classes, reducing latencies        caused by complex detection mechanisms and delay required to        ensure stable transition of the behavior of the motorized        prosthetic and/or orthotic device 200;    -   motorized prosthetic and/or orthotic device 200 joints limited        impedance in aerial phase increases ease of manipulation in        confined spaces and when maneuvering around obstacles;    -   cyclical locomotion tasks initiation is facilitated as the user        provides himself the proper pace and stride length; and    -   as the gait cycle patterns are not issued from a model, or        trajectory generation engine, or time-based mechanism, any        activity or gait phase may be interrupted at any instant without        compromising user support and safety,

Braking Implementation

The third class of lower-limb joint behavior, the Eccentric class, maybe advantageously addressed through a software-based braking mechanismimplementation. The Eccentric class of joint behavior is concerned withthe dissipation of energy by the joint of the motorized prosthetic ororthotic apparatus 140 (see FIG. 1). Generally speaking, the energy isinjected from an external source and requires dissipation in order toproperly manage joint behavior and resulting motion. For example, on thehuman healthy limb, Eccentric joint behavior is observed at eachextremity of the level walking swing phase, where it is required to stopthe knee joint motion due to shank inertia. Moreover, the use of theperturbation force matching behavior previously introduced in the aerialphase tends to accentuate this issue by giving a very low impedance tothe knee joint.

While multiple approaches exist to solve this type of problem, it isadvantageous to implement the Eccentric joint behavior class in areactive fashion to ensure constant behavior and performance from theuser standpoint. Moreover, it is advantageous to avoid the use of atrajectory-based mechanisms that only provide limited flexibility andrequire much tuning to account for inter-user variability.

Using the general framework provided by the variable gains impedancecontroller 50 shown in FIG. 5, explicit reactive behavior is added tothe basic formulation, leading to the block diagram of FIG. 7. From amotorized prosthetic and/or orthotic device 200 perspective,implementation of the Eccentric joint behavior class for the aerialmechanical lower-limb configuration 40 (see FIG. 4) is equivalent tomanaging braking using the actuator 144. Actuator 144 braking in theaerial configuration 40 may be achieved using many approaches: reductionof the perturbation force matching 64 (see FIG. 6) effort in order toallow natural dissipation to take place, increase of the actuator 144impedance, and reversal of the actuator 144 motion such that the motortorque is found in the opposite direction as its velocity.

FIG. 7 shows a variable gains impedance controller 70 based on thevariable gains impedance controller 50 of FIG. 5 to which a brakingprocess has been integrated, the braking process including all of theapproaches described above. The braking joint behavior, associated witheccentric muscle activation on the sound limb, is active at all time butits action is controlled through a set of logical conditions on varioussystem variables. Thus, the braking process may be integrated in thevariable impedance controller 70 as a braking feedback transfer function72 subject to conditional execution. Execution conditions are based onthree main variables: actuator 144 output position Θ, actuator 144output velocity {dot over (Θ)} and lower-limb mechanical configuration30, 40. While in the aerial configuration 40, the braking process isactivated when a velocity threshold is reached in the vicinity of, forexample, an end-of-motion bumper or a software-defined maximum targetflexion angle.

FIG. 8 illustrates the velocity-position subspace 80 in which thebraking process operates. The hatched regions 82, 84 represent theregions where the braking process is activated, otherwise, the brakingprocess remains passive, i.e. f(Θ,{dot over (Θ)})=0. In the simplestembodiment of the braking process, the regions 82, 84 in which thebraking process activates may be defined as both ends of the knee jointmotion range of the motorized prosthetic or orthotic apparatus 140 (seeFIG. 1). Whenever the joint enters one end of the motion range,delimited by the actuator 144 output angular position extension Θ_(ext)and output angular position flexion Θ_(f) activation thresholds, whileshowing velocity towards the nearest physical motion stop (i.e. bumper)superior to the output angular velocity activation threshold {dot over(Θ)}_(th), the braking process is activated in order to stop the jointand segment motion of the motorized prosthetic or orthotic apparatus 140before reaching the physical motion stop. Upon activation of the brakingprocess, the braking feedback transfer function 72 generates an outputsignal 73 that is removed from the net force command balance 74 that isused as the actuator 144 command signal.

More specifically, the braking feedback transfer function 72 may bedefined as:

$\begin{matrix}{{{{if}\mspace{14mu} ( {\overset{.}{\Theta} < {- {\overset{.}{\Theta}}_{th}}} )}\&\&{( {\Theta < \Theta_{ext}} )\text{:}}}{{f = \frac{\overset{.}{\Theta}}{( {\Theta + \Delta} )^{2}}};}{{{if}\mspace{14mu} ( {\overset{.}{\Theta} > {\overset{.}{\Theta}}_{th}} )}\&\&{( {\Theta > \Theta_{f}} )\text{:}}}{{f = \frac{\overset{.}{\Theta}}{( {\Theta_{\max} + \Delta - \Theta} )^{2}}};}} & {{Equation}\mspace{14mu} 3}\end{matrix}$

-   -   otherwise:

f=0;

-   -   where        -   f is the braking feedback transfer function;        -   Δ is the position offset;        -   Θ is the actuator output position measurement;        -   Θ_(ext) is the actuator output position extension activation            threshold;        -   Θ_(f) is the actuator output position flexion activation            threshold;        -   Θ_(max) is the actuator output maximum achievable position;        -   {dot over (Θ)} is the actuator output velocity measurement;            and        -   ±{dot over (Θ)}_(th) is the actuator output angular velocity            activation threshold.

Based on Equation 3, the braking feedback transfer function 72, orbraking force, may then be defined as the ratio of the joint velocity{dot over (Θ)} to the squared position measurement Θ, where an offset Δis added to ensure that the braking force remains a finite quantitywhile reaching the motion range end. Using such a relationship tocompute the braking force to be accounted in the net actuator 144command calculation allows the creation of a braking force thatincreases as the joint move towards the motion end while maintaining asignificant velocity, while not restricting motion in the directionopposite to the motion end. Such behavior differs from simply increasingthe joint impedance of a motion tracking control scheme, as the behaviorherein defined is characterized by its single sided action.

While Equation 3 is defined to ensure that braking occurs prior toreaching the hardware motion stops, it is also possible to dynamicallyconfigure the braking process parameters in order to modify the locationin the motion range where braking occurs. Hence, this braking processmay also be advantageously used in order to manage swing phase heel riseduring cyclical portions, or for other specialized functions such asmotion range limitations during rehabilitation or training processes.While the first suggested use could be fully automated throughdefinition of the proper detection and adjustment mechanism in theinference layer 120 (see FIG. 1), the second suggested use wouldoptimally be linked to a user/clinician interface device, allowing thisinterface device to configure the motorized prosthetic and/or orthoticdevice 200 according to the requirements of the rehabilitation/trainingprocess.

Referring back to FIG. 7, from a variable gains impedance controller 70standpoint, the additional behavior is integrated as a supplementaryfeedback term 73 that is added to the basic formulation. Referring backto the motorized knee prosthesis 10 of FIG. 2, the braking feedbacktransfer function 72 uses as input the measured relative positionbetween actuator 12 output and thigh segment (not shown), and theestimated joint velocity. Joint velocity may be estimated using an ARMA,i.e. Auto Regressive Moving Average process, which is shown to providean estimate of sufficient quality while minimizing the requirement forhardware sensors. Upon fulfillment of the conditions illustrated in FIG.8, the position measurement and velocity estimates are then used inorder to compute the amplitude of the braking force.

As previously discussed, the braking force then acts on the variablegains impedance controller 70 behavior by reducing the force feedbacksustaining the perturbation force matching process 64. Hence, thebraking force first compensates for the force feedback term 55, leavingthe actuator 144 in a passive mode. Leaving the actuator 144 in apassive mode when the joint is actually driven by Inertial forces allowsthe use of the motorized prosthetic or orthotic apparatus 140 poorpassive dynamics in order to fulfill the objective of the current jointreactive behavior, i.e. dissipation of energy in order to break jointmotion. If the use of passive braking is not sufficient to stop themotion, the form of the braking transfer function 72 defined by Equation3 generates a braking force that gains in amplitude as the jointcontinues to move towards the motion stop. As the braking force becomesgreater than the perturbation matching force term, i.e. force feedbackterm 55, the actuator 144 starts generating a force in the directionopposed to the motion, which results in a quick stop of the motion. Inthe swing phase, i.e. the aerial phase 40, the actuator 144 behaviordepends on the balance between the contribution of the force feedbackterm 55, and the proportional-derivative terms i.e. Θ and {dot over(Θ)}. Since K_(P) and K_(D) are set to 0 for the swing phase, actuator144 behavior is then defined by the sum of the force feedback term 55and the supplementary feedback term 73. Based on the definition of thebreaking transfer function 72, the force feedback term 55 is firstcancelled out by the supplementary feedback term 73 as the latterincreases. As the supplementary feedback term 73 becomes larger than theforce feedback term 55, the force following is effectively cancelled outand the supplementary feedback term 73 becomes the main contributor tothe amplitude and direction of the command signal sent to the actuator144. By their nature and definition, the force feedback term 55 and thesupplementary feedback term 73 will always be of opposite sign as thefirst one tries to follow the shank segment velocity while the secondones tries to control the shank segment velocity.

The above described braking process has been found to be very efficientand robust to inter-subjects variability as well as properly fulfillingdesired cyclical or non-cyclical locomotion tasks. Moreover, thereactive and self-adjusting nature of the braking process allows togreatly reduce dependency on locomotion portion, gait speed or userphysiological parameters, with respect to other types of systems relyingon position control. Such implementation of the Eccentric joint behaviorclass implicitly manages end-of-motion collisions in a way that is veryadaptable to various locomotion tasks and shows very high synergy withthe user due to its physiologically-compliant nature.

One indirect benefit associated with the use of such a braking processwith respect to other approaches based on hardware mechanisms arise fromthe fact that the actuator 144 is used in a regenerating mode.Regeneration occurs in an electrical motor when torque and velocity arein opposite directions. In such a case, assuming that proper driveelectronics are used, the motor starts acting as a generator and may beself-sufficient as far as power consumption is concerned. Implementationof the braking process herein defined then leads to a positive powerbalance, as mechanical work is generated without drawing any power fromthe power source of the motorized prosthetic and/or orthotic device 200.Furthermore, depending on the quantity of energy required to bedissipated using the braking process, i.e. depending on locomotiontasks, gait speed, user gait style and user physiological parameters, itmay also be possible to generate more energy than what is required bythe actuator's 144 motor to ensure satisfactory braking. Assuming that asuitable power supply architecture is used, for example the power supplydescribed in U.S. Pat. No. 7,230,352 entitled “COMPACT POWER SUPPLY” byBedard et al., it may then be possible to store the extra energy, whichis not required by the actuator 144 motor in order to sustain braking,for later use. From a motorized prosthetic and/or orthotic device 200perspective, this allows an increase in autonomy without any additionalcomponents.

Energy Injection Implementation

The fourth class of lower-limb joint behavior, the Concentric class, maybe advantageously addressed through an energy injection implementation.The Concentric class of joint behavior occurs whenever the lower-limbjoints of the motorized prosthetic and/or orthotic device 200 are usedin order to generate mechanical power or inject energy to sustainoverall gait. While some behaviors described above could be easilyimplemented on passive lower-limb prosthetic or orthotic joints,integration of a highly performing concentric behavior requires theavailability of mechanical power generation capabilities at the joint.While it might be argued that the use of simple passive mechanicalcomponents, for example springs, accumulators, etc., may allow energystoring and return, the limitations in power generation capabilitieswith respect to specific gait requirements make it difficult to achievesomething close to a reactive behavior using these passive mechanicalcomponents.

While obvious occurrence of Concentric joint behavior are found inlocomotion tasks such as stairs ascent, incline plane ascent orsit-to-stand transfer, the implementation of the Concentric reactivebehavior aims at fulfilling gait requirements different from the onesfound in these locomotion tasks. The concentric joint behaviorimplemented as reactive behavior is related to the implementation ofjoint motion in order to enforce sufficient toe clearance in bothcyclical and non-cyclical locomotion tasks.

Toe clearance management is an important feature of any motorizedprosthetic and/or orthotic device 200, as this feature may dramaticallyinfluence the overall device usability. While multiple approaches existregarding management of toe clearance on both passive and activelower-limb devices currently on the market, they all lack the ability toproperly manage toe clearance for both cyclical and non-cyclicallocomotion tasks, without affecting the device's usability or requiringthe user to adopt specific behaviors, often leading to a pathologicalgait.

From that respect, the definition of a generalized joint behavioraddressing the toe clearance management problem in a physiologicallycoherent and robust manner appears to be the most straightforwardsolution.

Concentric behavior targeting basic toe clearance management is thendefined as a low-level reactive behavior allowing to connect sensoryinput from the sensors 142 to a pre-defined joint behavior. Upondetection of the motorized prosthetic and/or orthotic device 200transition from the interacting 30 to the non-interacting 40 mechanicalconfiguration (see FIGS. 3 and 4, respectively), energy injection at thejoint level is triggered and takes place as a supplementary feed forwardterm in the formulation of the variable gains impedance controller's 70of FIG. 7.

Since the requirements for any Concentric joint action targeting toeclearance are both user-specific and locomotion task specific, energyInjection is advantageously implemented in conjunction with auser-interface device allowing the customization of the basic energyinjection implementation's behavior. Through the combination of theenergy injection implementation and associated user-interface device, itmay be possible to define a general baseline behavior. In order toaccount for more complex concentric joint behavior requirements, it maybe possible to couple this general baseline behavior with higher levelinference engines that will allow the dynamic modification of the energyinjection amplitude, timing and duration. Such modifications depend onthe nature of the task currently performed by the user.

From an inference layer 120 perspective (see FIG. 1), threespecializations are considered to affect the general baseline behaviorof the energy injection Implementation and are associated withsustaining adequate toe clearance and heel rise in level walking, stairsor incline ascent, and stairs or incline descent. Appropriate adjustmentof the energy injection implementation reactive layer 130 parameters bythe inference layer 120 engines may ensure fulfillment of these threespecializations in a seamless manner.

Referring to FIG. 9, from a variable gains impedance controller 90standpoint, the energy injection implementation may be advantageouslyimplemented as a feed forward transfer function 92 acting as a discretepulse generator which directly injects a force pulse 93 at the output 94of the positional terms of the variable impedance controller 90 upontriggering of the transfer function discrete input Γ. As discussedabove, the triggering mechanism may consist in a low-level detection ofthe transition from the interacting configuration 30, i.e. foot incontact with the ground, to the non-interacting configuration 40, i.e.aerial lower-limb configuration. The feed forward transfer function 92,g(s), may take a wide variety of form, for example a pulse typewaveform. Other types of discrete-time waveforms may also be defined forthat specific purpose (e.g., saw tooth, exponential, etc.).

Hence, upon transition to the non-interacting configuration 40, both theenergy injection and perturbation force matching 64 (see FIG. 6) areactive, ensuring that minimal motorized prosthetic or orthotic apparatus140 joint flexion take places before the joint is left in its minimalimpedance state. While this last sequence of event takes place withoutconsideration of the cyclical nature of the task being executed, morespecific actions are expected to take place and sustain the completeswing phase of cyclical locomotion tasks, such that proper footclearance and subsequent foot placement takes place.

While the benefits associated with the behavior described above for thecyclical locomotion tasks are quite straightforward, it is thecapability to properly manage requirements associated with non-cyclicaltasks that make the implementation of the concentric joint behaviorinteresting for a motorized prosthetic and/or orthotic device 200.Combination of the Concentric behavior allowing the enforcement of basictoe clearance in limited ambulation tasks to the Isometric behaviorallowing support in the absence of motion during the contact phasewithout consideration of the knee flexion angle at which the groundcontact occurs greatly eases the burden associated with the manipulationof a lower-limb motorized prosthetic and/or orthotic device 200 withrespect to more conventional designs.

Moreover, it was shown in experimental testing that the combination ofthe energy injection implementation with the force matching and forcerejection implementations greatly enhance the usability of the motorizedprosthetic and/or orthotic device 200 when facing constrainedenvironments, obstacles, or other types of situations that cannot becharacterized through typical locomotion tasks, Enforcement of a certainknee flexion angle through the effects of the energy injectionimplementation also facilitates the implementation of less pathologicalgait habits in limited ambulation, as stance phase knee flexion iseasily obtained and provide adequate support, without being overlystiff. Hence, improved physiological interaction between the user andits motorized prosthetic and/or orthotic device 200 may be obtained.

It is to be understood that the force matching and force rejectionimplementations, the braking implementation and the force injectionimplementation may be integrated individually or in any combinationthereof into a conventional variable gains impedance controller to forma reactive layer control system for orthotic or prosthetic devices.

Although the present invention has been described by way of particularnon-limiting illustrative embodiments and examples thereof, it should benoted that it will be apparent to persons skilled in the art thatmodifications may be applied to the present particular embodimentwithout departing from the scope of the present invention.

REFERENCES

-   [1] Hogan, N., Impedance Control: An Approach to Manipulation: Part    I—Theory. ASME Journal of Dynamic Systems, Measurement and Controls,    vol. 107, pp. 1-7, 1985.-   [2] Hogan, N., Impedance Control: An Approach to Manipulation part    II—Implementation, ASME Journal of Dynamic Systems, Measurement and    Controls, vol. 107, pp. 8-16, 1985.

[3] Hogan, N., Impedance Control: An Approach to Manipulation: PartIII—Applications, ASME Journal of Dynamic Systems, Measurement andControls, vol. 107, pp. 17-24, 1985.

1-84. (canceled)
 85. A system for controlling an apparatus, the systemcomprising: a force sensor; a torque sensor; and an impedance controllerconfigured to control a motorized prosthetic or orthotic apparatus, aprocessor in communication with the impedance controller, the forcesensor, and the torque sensor, the processor configured to: receive afirst signal from the force sensor, the first signal indicative of aninteraction between the apparatus and ground; receive a second signalfrom the torque sensor, the second signal indicative of torque at ajoint; and determine a result by multiplying the second signal by avalue, wherein the result is based on the first signal.
 86. The systemof claim 85, wherein the value is a positive value and substantiallyunitary when the first signal indicates an interaction, wherein thevalue is a negative value and substantially unitary when the firstsignal indicates an absence of interaction.
 87. The system of claim 85,further comprising a position sensor, wherein the processor is furtherconfigured to: receive a third signal from the position sensor, thethird signal indicative of position of the joint; and responsive to thefirst signal indicating an interaction, the processor is furtherconfigured to: multiply the second signal by a positive value; estimatea velocity of the joint from the third signal; multiply a differencebetween a velocity set-point and the estimated velocity by a velocitygain value; and multiply a difference between a position set-point andthe third signal by a position gain value.
 88. The system of claim 87,wherein the processor is further configured to: provide a brakingfeedback value to the impedance controller based at least in part on adetermination that the velocity is greater than a first threshold andthe third signal indicative of position is greater than a secondthreshold, wherein the braking feedback value increases joint resistanceto motion.
 89. The system of claim 88, wherein the braking feedbackvalue is based on a ratio of the velocity to a square of an indicatedposition offset with a motion stop target of the apparatus.
 90. Thesystem of claim 85, wherein the processor is further configured toprovide a pulse to impedance controller based on a determination thatthe force sensor indicates a transition from an interaction between theapparatus and the ground to an absence of interaction between theapparatus and the ground.
 91. The system of claim 90, wherein theprocessor is further configured to dynamically adjust characteristics ofthe pulse, wherein the dynamically adjusted characteristics of the pulseinclude a combination of one or more of an amplitude, a duration, or atiming.
 92. The system of claim 85, wherein the first signal is ameasure of load at a shank portion of the motorized prosthetic ororthotic apparatus, and wherein the interaction between the apparatusand the ground is determined by a single value hysteretic thresholdingof the first signal received from the force sensor.
 93. The system ofclaim 85, wherein the system further comprises a position sensor, andwherein the processor is further configured to: receive a third signalindicative of a position of the joint from the position sensor; andresponsive to a determination that the first signal indicates an absenceof interaction, the processor is further configured to: multiply thesecond signal by a negative value and providing the result to theimpedance controller, estimate a velocity of the joint from the thirdsignal, multiply a difference between a velocity set-point and theestimated velocity by a velocity gain value and providing the result tothe impedance controller, and multiplying the difference between aposition set-point and the third signal by a position gain value andproviding the result to the impedance controller.
 94. The system ofclaim 85, wherein the processor is further configured to: receive asignal indicative of velocity of the joint; receive a signal indicativeof position of the joint; and responsive to a determination that thefirst signal indicates an absence of interaction, the processor isfurther configured: multiply the second signal by a negative value andproviding the result to the impedance controller, multiply a differencebetween a velocity set-point and the signal indicative of velocity by avelocity gain value and providing the result to the impedancecontroller, and multiply a difference between a position set-point andthe signal indicative of position by a position gain value and providingthe result to the impedance controller.
 95. A system for controlling anapparatus, the system comprising: an impedance controller configured tocontrol an actuator of an apparatus, wherein the impedance controller isfurther configured to: receive a first signal indicative of a torque ata joint of the apparatus; determine a torque feedback based at least inpart on the first signal and a gain factor, wherein during swing phasethe gain factor is a first gain factor, wherein during stance phase thegain factor is a second gain factor that is different from the firstgain factor; and output a control signal to the actuator based at leastin part on the torque feedback, wherein the actuator adjusts jointresistance to motion based at least in part on the control signal. 96.The system of claim 95, wherein the apparatus is a lower-limb prostheticdevice or a lower-limb orthotic device.
 97. The system of claim 95,wherein the actuator is coupled to a prosthetic limb member, and theactuator forms at least a portion of the joint of the apparatus.
 98. Thesystem of claim 95, the actuator is coupled to an orthotic limb member,and the actuator forms at least a portion of the joint of the apparatus.99. The system of claim 95, wherein one of the first gain factor or thesecond gain factor is a positive value and the other one of the firstgain factor or the second gain factor is a negative value.
 100. Thesystem of claim 95, wherein the impedance controller is furtherconfigured to receive a second signal indicative of a position of thejoint, wherein the control signal is further based at least in part onthe second signal.
 101. The system of claim 95, further comprising atorque sensor, wherein the torque sensor produces the first signal. 102.The system of claim 95, wherein the impedance controller is furtherconfigured to receive a second signal indicative of an interactionbetween the apparatus and a walking surface.
 103. The system of claim102, wherein the control signal is further based at least in part on thesecond signal.
 104. The system of claim 102, wherein the impedancecontroller is further configured to control the actuator to reduce jointflexion based at least in part on a determination that the second signalindicates an absence of interaction between the apparatus and thewalking surface.